Assessment of the lung parenchyma by means of magnetic resonance images

ABSTRACT

A method for the assessment of the lung parenchyma in a human or an animal is indicated using a series of magnetic resonance (MR) images of the lung parenchyma acquired at different breathing positions in the same human or animal. The method comprises at least the steps of a.) estimating a change of the lung volume V L  between the different breathing positions, b.) determining a signal intensity SI({right arrow over (x)} i ) in at least one same region or position {right arrow over (x)} i  of the lung parenchyma for each of the MR images, and c.) determining at least one respiratory index α({right arrow over (x)} i ) according to the formula 
     
       
         
           
             
               
                 
                   
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TECHNICHAL FIELD

The present invention concerns a method for the assessment of the lungparenchyma in a human or an animal using magnetic resonance (MR) images.The present invention also concerns a computer program comprisingexecutable instructions for the assessment of the lung parenchyma in ahuman or an animal by means of MR images as well as a MR imaging systemconfigured for acquiring MR images of a human or an animal lung and forassessing the lung parenchyma based on these images.

PRIOR ART

Periodical bidirectional airflow is the primary process taking place inthe lung, enabling gas exchange in the alveoli. As a result, themeasurement of both, global and local lung ventilation is a keyparameter not only to diagnose and understand pathophysiology in lungdiseases, but also for monitoring the progression and the response totherapeutic procedures.

A global measure of lung ventilation can be obtained by conventionalwhole-lung spirometry, plethysmography or multiple-breath washout tests.These conventional pulmonary function tests, however, fail to provideinformation about regional lung processes which can provide fundamentalinsights into pathophysiologic mechanisms or improve therapies by atailoring for specific lung regions and disease phenotypes.

A broad spectrum of imaging modalities is available to assess theregional lung ventilation. Single-photon emission computed tomography(SPECT), planar scintigraphy and positron emission tomography (PET) arenuclear medicine imaging modalities based on radioactive labeled traceradministration. Despite their relatively poor spatial and temporalresolution, SPECT and PET are commonly used in the clinical routine forlung functional imaging. High spatial resolution ventilation imaging inshort scanning times is possible by means of dynamic contrast-enhancedcomputed tomography (CT).

CT-methods do derive ventilation-related information based ontwo-dimensional (2D) images and models tailored for CT are presented bySimon B A: Non-Invasive Imaging of Regional Lung Function using X-RayComputed Tomography. J Clin Monit 2000; 16: 433-442 and by Guerrero T,Sanders K, Castillo E, Zhang Y, Bidaut L, Pan T, Komaki R: Dynamicventilation imaging from four-dimensional computed tomography. Phys.Med. Biol. 51 (2006) 777-791. These models, when further modified, canbe used to derive a measure of the lung emphysema progression thanks toa density-volume slope model (sponge model) between baseline andfollow-up examinations, as presented by Staring M, Bakker M E, ShamoninD P, Stolk J, Reiber J H C, Stoel B C: Towards local estimation ofemphysema progression using image registration. Proc SPIE2009:72590O-72590O-9. A density-volume slope model (sponge model) for CTis proposed in this document to correct for the different inspiratoryvolumes between baseline and follow-up examinations, which would flawthe emphysema progression measurement.

Despite the constant progress of nuclear medicine and CT in reducingradiation doses, exposure to any ionizing radiation is still of concern,especially for children, pregnant women, or patients who requirefrequent follow-up examinations.

Alternatively, magnetic resonance imaging (MRI) is an attractiveradiation-free imaging modality, and over the last decade a variety ofpowerful methods have been developed for the evaluation of both lungmorphology and function. Among them, non-proton based lung imaging withhyperpolarized noble gases, such as Helium-3 or Xenon-129, hasdemonstrated the ability to measure various pulmonary functionalbiomarkers; furthermore, the pulmonary ventilation can also bevisualized in MRI using inert fluorinated gas tracers which has shownlately high quality images similar to those from hyperpolarized gases.Gas ventilation MRI, however, is limited to few specialized medicalcenters, mainly due to high cost, complexity and requirements forexperimental hardware.

In contrast, non-invasive, proton-based, imaging of regional lungfunction is possible in vivo using a technique known as Fourierdecomposition (FD) MRI (Bauman G, Puderbach M, Deimling M, et al.:Non-contrast-enhanced perfusion and ventilation assessment of the humanlung by means of Fourier decomposition in proton MRI. Magn Reson Med2009; 62:656-664). This approach utilizes dynamic balanced steady statefree precession (SSFP) imaging in free-breathing with subsequent imageregistration and voxel-wise spectral analysis to extract localventilation and perfusion information. FD MRI is based on nativebalanced SSFP imaging, thus requires no administration of contrastagent, but from the required time-resolution to separate perfusion fromventilation, FD MRI is a single slice method. As a result, a stack ofcoronal 2D slices needs to be acquired to cover the whole lung volume,thus proffering relatively low through plane resolution.

By means of a MRI method referred to as 3D ultra-fast balanced SSFP(ufSSFP), it is possible to acquire three-dimensional (3D), i.e.volumetric chest scans of the lung at different inspiratory volumes(multi-volume 3D MRI) (Bieri O: Ultra-fast steady state free precessionand its application to in vivo 1H morphological and functional lungimaging at 1.5 tesla. Magn Reson Med 2013; 70:657-663).

It could be shown, that 3D ufSSFP is sensitive to depict the signalintensity modulations of the lung parenchyma arising by diverseinspiratory levels (Pusterla O, Bauman G, Bieri O: How volume affectsthe pulmonary MRI signal: Investigations with 3D ultra-fast balancedSteady-State Free Precession. Edited by Proc. Intl. Soc. Mag. Reson.Med. 23., 2015:1481). In the method introduced by this document, thelung parenchyma is mechanically considered as a sponge-like tissue,which periodically expands and contracts following the movement of thediaphragm and chest walls, resulting in bidirectional air-flow andventilation.

SUMMARY OF THE INVENTION

It is an object of the present invention to provide a proton-based andcontrast-agent-free MRI method which provides reliable regionalinformation related to pulmonary respiration. The method should beapplicable in the clinical setting and particularly also in patientswith shortness of breath and cooperative children.

In order to achieve this object, the present invention provides a methodfor the assessment of the lung parenchyma in a human or an animal usinga series of magnetic resonance (MR) images, in particular of 3D MRimages, of the lung parenchyma acquired at different breathing positions(thus 4D-MRI in case of 3D MR images) in the same human or animal,comprising at least the following steps:

-   -   a.) estimation of a change of the lung volume V_(L) (the whole        lung volume or a portion of the whole lung volume) between the        different breathing positions;    -   b.) determination of a signal intensity SI({right arrow over        (x)}_(i)) in at least one same region or position {right arrow        over (x)}_(i) of the lung parenchyma for each of the MR images;        and    -   c.) determination of at least one respiratory index α({right        arrow over (x)}_(i)) according to the formula

${\alpha \left( {\overset{\rightarrow}{x}}_{i} \right)} = {- {\frac{d\left( {\log \left( {{SI}\left( {\overset{\rightarrow}{x}}_{i} \right)} \right)} \right)}{d\left( {\log \left( V_{L} \right)} \right)}.}}$

Basically, the lung parenchyma can mechanically be considered as asponge-like tissue, which periodically expands and contracts thanks tothe movement of the diaphragm and chest walls, thus allowingbidirectional air-flow and ventilation. In a simple dry sponge modelwhere the sponge mass is constant, the signal intensity of MR imagesacquired for example by 3D ufSSFP is inversely proportional to the lungvolume V_(L):

${{{SI}\left( V_{L} \right)} = {\frac{\Lambda}{V_{L}} + \eta}},$

where Λ is a signal intensity scaling factor (parenchymal mass), and ηis the SI in the limit of infinite volume, i.e. noise. The assumption ofmass preservation in the dry sponge model, however, is not preciseenough to model the lung mechanics due to physiological changes such asthe blood volume variability during breathing and in presence ofventilation inhomogeneities or regional abnormalities. As a result, theequation of the dry sponge model needs to be modified to take thefollowing form of the adapted sponge model:

SI(V _(L))=Λ*V _(L) ^(−α)+η,

where α is referred to as the respiratory index, which maps lungexpansion and changes in pulmonary blood volume during breathing. On alogarithmic scale, the adapted sponge model yields

log(SI)=log(Λ*V _(L) ^(−α)+η)≈log(Λ*V _(L) ^(−α))=log(Λ)−α*log(V _(L)),

where η<<Λ*V_(L) ^(−α) is assumed, and α can be estimated globally(considering the mean SI over the whole lung volume, SI_(global)) using

$\alpha_{global} = {- {\frac{d\left( {\log \left( {SI}_{global} \right)} \right)}{d\left( {\log \left( V_{L} \right)} \right)}.}}$

Similarly, after spatially matching the lung structures of themulti-volume acquisitions, for example by a deformable registrationalgorithm, the respiratory index can be estimated locally at theposition(s) {right arrow over (x)}_(i) (e.g. voxel wise), in order toyield the formula according to step c.):

${\alpha \left( {\overset{\rightarrow}{x}}_{i} \right)} = {- {\frac{d\left( {\log \left( {{SI}\left( {\overset{\rightarrow}{x}}_{i} \right)} \right)} \right)}{d\left( {\log \left( V_{L} \right)} \right)}.}}$

The global respiratory index α_(global) is a measure for blood masschange during the respiratory cycle, as a function of the total lungvolume:

-   -   α_(global)=1: no mass is leaving nor entering the system, and        the dry sponge model is satisfied.    -   α_(global)>1: blood leaves the lung and its density decreases        faster than in the dry sponge model.    -   α_(global)<1: blood enters the lung and its density decreases        slower than in the dry sponge model.

The local respiratory index α({right arrow over (x)}_(i)) is a measurefor the local change of the SI (e.g. in a certain lung region orposition, such as in a voxel) over the respiratory cycle, as a functionof the total lung volume:

-   -   For α({right arrow over (x)}_(i))=1, the SI decrease is as in        the dry sponge model: The voxel expansion is proportional to the        lung-average expansion and no blood is entering or leaving the        voxel or the voxels of the respective region or position {right        arrow over (x)}_(i).    -   For α({right arrow over (x)}_(i))<1 during inspiration, the SI        of the respective region or position {right arrow over (x)}_(i)        decreases slower than can be contributed to the increase of the        total lung volume, meaning that pulmonary blood is entering the        system or a region with a reduced local ventilation which        expands at a rate lower than the entire lung-average is present.    -   For α({right arrow over (x)}_(i))>1 the opposite is taking        place: Blood is leaving or there is overextended ventilation in        the respective region or position {right arrow over (x)}_(i)        (the expansion rate is higher than the entire lung average        expansion).

Thus, the local respiratory index α({right arrow over (x)}_(i)) asdetermined in step c.) represents a regional isotropic three-dimensional(3D) measure that directly relates to the local pulmonary function.

Thus, MR images that have been acquired beforehand, i.e. before stepb.), usually also before step a.), are used to determine at least onerespiratory index α({right arrow over (x)}_(i)). The acquisition of theMR images can constitute a further step a0.) of the method. The seriesof MR images comprises at least two images or preferably more, at leastfive, more preferably at least nine MR images, and most preferably atleast 13 MR images, of the lung parenchyma which are each acquired at adifferent breathing position of the same human or animal.

In step a0.), the series of MR images is preferably acquired by means ofa balanced steady state free precession (SSFP) imaging sequence, inparticular a 3D ufSSFP sequence as disclosed in Bieri O: Ultra-faststeady state free precession and its application to in vivo 1Hmorphological and functional lung imaging at 1.5 tesla. Magn Reson Med2013; 70:657-663, the disclosure of which is hereby incorporated byreference. A good through plane resolution can be achieved by applying a3D data acquisition sequence. Indubitably other pulse sequences, e.g.GRE, UTE, ZTE, in 2D or more space-time dimensions can be used asimaging technique.

For acquiring the MR images, a commercial MRI whole-bode scanner isusually used preferably having a main magnetic field of 1.5 T. Ofcourse, it is also possible to acquire the MR images on a 3 T scanner oron a MR scanner having a smaller or a higher main magnetic field.

Step a.) can be carried out before step b.), after step b.) or togetherwith step b.). Step c.) is usually carried out after completion of bothof steps a.) and b.).

For the determination of the signal intensities SI({right arrow over(x)}_(i)) in step b.), usually signal magnitudes are identified in theMR images at certain positions of the lung parenchyma. It is alsopossible to determine the signal intensities SI({right arrow over(x)}_(i)) in step b.) by averaging the signal magnitudes of the MRimages over certain regions of the lung parenchyma, in order to improvethe reliability of the respiratory index α({right arrow over (x)}_(i)).Thus, the MR images which are used to determine the signal intensitiesSI({right arrow over (x)}_(i)) represent the acquired MR data in imagespace (as opposed to the respective k-space data).

It is possible for the respiratory index α({right arrow over (x)}_(i))to represent a global measure of the function of the entire lung. Insuch a case, the region {right arrow over (x)}_(i) denotes the entirelung parenchyma and the obtained respiratory index α({right arrow over(x)}_(i))=α_(global) represents the function of the entire lung. Morepreferred, however, is an embodiment in which {right arrow over (x)}_(i)denotes a position or region that only represents a part of the entirelung, more preferably a part of the right or the left lung, mostpreferably a part of one of the pulmonary lobes. In this case, therespiratory index α({right arrow over (x)}_(i)) represents the functionof only the respective part of the lung. In a particularly preferredembodiment, a plurality of respiratory indices α({right arrow over(x)}_(i)) is determined, in order to assess the function of the lungparenchyma at a plurality of respective regions or positions {rightarrow over (x)}_(i) of the lung. Thus, in this case in which a pluralityof respiratory indices α({right arrow over (x)}_(i)) is determined, theindex i runs from 1 to n, i.e. i=1 . . . n, with n≧2, preferably n≧32,more preferably n≧64.

If steps a.), b.) and c.) are carried out for a plurality of regions orpositions {right arrow over (x)}_(i) of the lung parenchyma, then thedetermined respiratory indices α({right arrow over (x)}_(i)) arepreferably used to generate a two-dimensional or three-dimensional mapreflecting the regional function of the lung. The map is preferably acolor-coded map in which the color for each position or region {rightarrow over (x)}_(i) is dependent on the respiratory indices α({rightarrow over (x)}_(i)) as determined in step c.).

Preferably, the acquired MR images are spatially co-registered in thearea of the lung parenchyma, advantageously in the area of the entirelung parenchyma, and, in step b.), the signal intensity SI({right arrowover (x)}_(i)) is determined for each of the spatially co-registered MRimages. Thus, the positions or regions {right arrow over (x)}_(i) in theplurality of co-registered MR images directly correspond to the samerespective positions or regions {right arrow over (x)}_(i) of the lungparenchyma. The spatial co-registration of the MR images isadvantageously carried out for example by means of a deformable B-splinemass preserving image registration algorithm. Of course otherregistration algorithms or correlations analysis can be used.

Advantageously, the spatially co-registered MR images can bemedian-filtered, and, in step b.), the signal intensity SI({right arrowover (x)}_(i)) is determined for each of the spatially co-registered andmedian-filtered MR images. Median filtering is used to remove theoverlying vascular structure from the MR images and get access to theparenchymal signal. Other filtering and post-processing methodology canalso be advantageous.

A variety of possibilities are conceivable to estimate the change of thelung volume V_(L) (or a part of the lung volume, e.g. a segment or lobe)between the different breathing positions in step a.), such as forexample by means of a simple balloon. Spirometry measurements can alsobe applied for this purpose. In a preferred embodiment, however, theestimation of the lung volume V_(L) for each of the different breathingpositions is based on the acquired series of MR images or other imagingmodalities (e.g. ultrasounds). For doing so, each of the acquired MRimages can be segmented with respect to the lung parenchyma, such that,in step a.), the lung volume V_(L) or a portion of it can be estimatedfor each of the different breathing positions based on the segmented MRimages. An alternative method to estimate the change of the lung volumeis based on a calculation of the determinant of the spatial Jacobian ofthe transformation field T, which links the positions of the tissueregions in the different images: V({right arrow over (x)}_(i))=J({rightarrow over (x)}_(i))=det(∂T/∂{right arrow over (x)}_(i)) which can forexample be calculated voxel-wise or in regions {right arrow over(x)}_(i).

For the estimation of the lung volumes V_(L), the acquired MR images areadvantageously median-filtered prior to the segmentation, in order tofacilitate the segmentation.

The series of MR images preferably represent the lung parenchyma atregularly distributed breathing positions from full expiration tomaximum inspiration. In case the series of MR images only comprises twoimages, the images need to be at significantly different breathingposition (phase). For example, one of these two images can be acquiredat full expiration and the other at maximum inspiration. Thedetermination of α({right arrow over (x)}_(i)) in step c.) with twoimages only is straightforward: The difference between the signalintensities SI({right arrow over (x)}_(i)), divided by the difference ofthe respective lung volumes V_(L), can be calculated.

Preferably, however, in order to obtain a more reliable result, theseries of MR images comprises at least three images, more preferably atleast five images, even more preferably at least nine images, and mostpreferably at least 13 images. In this case, in step c.), the at leastone respiratory index α({right arrow over (x)}_(i)) is advantageouslydetermined by means of a linear fit. Otherwise, the respiratory indicesα({right arrow over (x)}_(i)) can be calculated between each two images,thus calculating α({right arrow over (x)}_(i), I_(p1), I_(p2)), whereI_(p1) and I_(p2) indicate two images at different breathing positions.

In order to enable a more precise and in particularly quantitativedetermination of the at least one respiratory index α({right arrow over(x)}_(i)) in step c.), a multiplication with the factor

$\frac{1}{1 - {S\; N\; R^{- 1}}}$

is further carried out, in order to yield the formula

${{\alpha \left( {\overset{\rightarrow}{x}}_{i} \right)} = {{- \frac{1}{1 - {S\; N\; R^{- 1}}}} \cdot \frac{d\left( {\log \left( {{SI}\left( {\overset{\rightarrow}{x}}_{i} \right)} \right)} \right)}{d\left( {\log \left( V_{L} \right)} \right)}}},$

wherein SNR denotes an estimated signal-to-noise ratio SI({right arrowover (x)}_(i))/η in which η represents the estimated intensity of noisein the MR images in the at least one region or position {right arrowover (x)}_(i).

The equation of the adapted sponge model, SI(V_(L))=Λ*V_(L) ^(−α)+η, canbe rewritten on a logarithmic scale to yield

${{\log ({SI})} = {\log \left( {\Lambda*{V_{L}^{- \alpha}\left( {1 + \frac{\eta}{\Lambda*V_{L}^{- \alpha}}} \right)}} \right)}},$

from which the respiratory index α can be estimated as follows:

${\frac{\partial{SI}}{\partial\left( {\log \left( V_{L} \right)} \right)} = {- \alpha}}{{\cdot \frac{\Lambda \cdot V_{L}^{- \alpha}}{{\Lambda \cdot V_{L}^{- \alpha}} + \eta}} = {{{- \alpha} \cdot \left( {1 - \frac{\eta}{{\Lambda \cdot V_{L}^{- \alpha}} + \eta}} \right)} = {\left. {{- \alpha} \cdot \left( {1 - \frac{\eta}{SI}} \right)}\mspace{20mu}\Rightarrow\alpha \right.:={{{- \frac{1}{1 - \frac{\eta}{\overset{\_}{SI}}}} \cdot \frac{d\left( {\log ({SI})} \right)}{d\left( {\log \left( V_{L} \right)} \right)}} = {{- \frac{1}{1 - {S\; N\; R^{- 1}}}} \cdot \frac{d\left( {\log ({SI})} \right)}{d\left( {\log \left( V_{L} \right)} \right)}}}}}}$

Generally, the noise η might vary for each patient, for each MRIscanner, and for each acquisition, and becomes prominent particularly inpulmonary MRI. This is caused by the physical characteristic of thelung, e.g. low and variable density together with susceptibilityeffects, which cause both a low detectable SI and strong SI variationsamongst the pulmonary tissue (e.g. dorsal/ventral part,dependent/non-dependent, perfused/non-perfused). As a result, the factor

$\frac{1}{1 - {S\; N\; R^{- 1}}}$

plays a crucial role in determining the precise value of α. By applyingthis factor, not only a qualitative assessment, but a quantitativeexamination of the regional lung function becomes possible, which inturn enables follow-up studies, e.g. of obstructive pulmonary diseasepatients.

Additionally, a magnetic resonance imaging (MRI) system is providedwhich is configured to carry out the method as indicated. Usually, sucha MRI system comprises at least:

-   -   a main magnet for generating a main, particularly static        magnetic field B₀ at a location of a sample, i.e. a human or an        animal, to be imaged, in order to at least partly align nuclear        spins of the sample;    -   an excitation module for applying a sequence of radio frequency        (RF) pulses to the sample, in order to repeatedly excite the        nuclear spins of the sample;    -   a gradient module for generating temporary magnetic gradient        fields at a location of the sample;    -   an acquisition module, in particular an image acquisition        module, for acquiring the magnetic resonance (MR) signals        produced by excited nuclear spins of the sample; and    -   a control module configured for controlling the excitation        module, the gradient module and the acquisition module such,        that a series of MR images of the lung parenchyma of a human or        an animal can be acquired at different breathing positions.

Furthermore, the MRI system comprises an analysis module beingconfigured for determining a signal intensity SI({right arrow over(x)}_(i)) in at least one same region or position {right arrow over(x)}_(i) of the lung parenchyma for each of the MR images and beingconfigured for determining at least one respiratory index α({right arrowover (x)}_(i)) according to the formula

${{\alpha \left( {\overset{\rightarrow}{x}}_{i} \right)} = {- \frac{d\left( {\log \left( {{SI}\left( {\overset{\rightarrow}{x}}_{i} \right)} \right)} \right)}{d\left( {\log \left( V_{L} \right)} \right)}}},$

wherein V_(L) denotes a lung volume for each of the different breathingpositions.

It is preferred, that the magnetic field B₀ generated by the main magnetis essentially uniform at least in the region of the sample. Themagnitude of the magnetic field generated by the main magnet ispreferably larger than 0.5 Tesla, more preferably larger than 1 Tesla,even more preferably larger than 2 Tesla and most preferably larger than5 Tesla. With a stronger magnetic field, a better signal-to-noise ratiocan be achieved, but imaging artifacts are also more pronounced.

The MRI system preferably also comprises an image reconstruction modulefor reconstructing images based on the acquired k-space data.

Preferably, the analysis module is further configured for estimatingchange of the lung volume V_(L) (or a portion of it) between thedifferent breathing positions based on the MR images. For this purpose,the analysis module can particularly be configured to median filter andsegment the MR images, wherein the segmentation process isadvantageously assisted by user interaction.

In a particularly preferred embodiment, the analysis module is alsoconfigured for carrying out a multiplication with the factor

$\frac{1}{1 - {S\; N\; R^{- 1}}}$

in the determination of the at least one respiratory index α({rightarrow over (x)}_(i)), in order to yield the formula

${{\alpha \left( {\overset{\rightarrow}{x}}_{i} \right)} = {{- \frac{1}{1 - {S\; N\; R^{- 1}}}} \cdot \frac{d\left( {\log \left( {{SI}\left( {\overset{\rightarrow}{x}}_{i} \right)} \right)} \right)}{d\left( {\log \left( V_{L} \right)} \right)}}},$

wherein SNR denotes an estimated signal-to-noise ratio SI({right arrowover (x)}_(i))/η in which η represents the estimated intensity of noisein the MR images in the at least one region or position {right arrowover (x)}_(i).

Furthermore, a computer program, preferably stored on a storage devicereadable by a computer, is provided comprising executable instructionsto carry out the method as indicated when being executed e.g. in aprocessor of a computer, in particular in a processor of a MRI system orin a processer being connected with a MRI system. The computer programis usually realized as a computer program code element which comprisescomputer-implemented instructions to cause a processor to carry out aparticular method. It can be provided in any suitable form, includingsource code or object code. In particular, it can be stored on acomputer-readable medium or embodied in a data stream. The data streammay be accessible through a network such as the Internet.

Thus, the computer program serves for the assessment of the lungparenchyma in a human or an animal by means of a series of magneticresonance (MR) images of the lung parenchyma acquired at differentbreathing positions in the same human or animal. The computer programcomprises at least executable instructions to:

-   -   b.) determine a signal intensity SI({right arrow over (x)}_(i))        in at least one same region or position {right arrow over (x₁)}        of the lung parenchyma for each of the MR images; and    -   c.) determine at least one respiratory index α({right arrow over        (x)}_(i)) according to the formula

${{\alpha \left( {\overset{\rightarrow}{x}}_{i} \right)} = {- \frac{d\left( {\log \left( {{SI}\left( {\overset{\rightarrow}{x}}_{i} \right)} \right)} \right)}{d\left( {\log \left( V_{L} \right)} \right)}}},$

wherein V_(L) denotes a lung volume for each of the different breathingpositions.

The computer program can also comprise executable instructions forcontrolling a magnetic resonance imaging (MRI) system as described. Inthis case, the computer program also comprises executable instructionsto employ a MR imaging sequence in an additional step a0.), such as a 3DufSSFP imaging sequence, on the MRI system in such a way, that a seriesof MR images of the lung parenchyma of a human or an animal is acquiredat different breathing positions.

The computer program advantageously additionally comprises executableinstructions to estimate, in an additional step a.), the change of thelung volume V_(L) between the different breathing positions based on theMR images, in particular according to one of the methods as set outabove in respect to this purpose.

The executable instructions of the computer program to determine atleast one respiratory index α({right arrow over (x)}_(i)) in step c.)preferably also comprise a multiplication with the factor

$\frac{1}{1 - {SNR}^{- 1}},$

in order to yield the formula

${{\alpha \left( {\overset{\rightarrow}{x}}_{i} \right)} = {{- \frac{1}{1 - {SNR}^{- 1}}} \cdot \frac{d\left( {\log \left( {{SI}\left( {\overset{\rightarrow}{x}}_{i} \right)} \right)} \right)}{d\left( {\log \left( V_{L} \right)} \right)}}},$

wherein SNR denotes an estimated signal-to-noise ratio SI({right arrowover (x)}_(i))/η in which η represents the estimated intensity of noisein the MR images in the at least one region or position {right arrowover (x)}_(i).

SHORT DESCRIPTION OF THE FIGURES

Preferred embodiments of the invention are described in the followingwith reference to the drawings, which only serve for illustrationpurposes, but have no limiting effects. In the drawings it is shown:

FIG. 1 a schematic illustration of a MRI system for carrying out theinventive method according to an inventive embodiment;

FIG. 2 a flow chart illustrating a preferred example for carrying outthe inventive method;

FIG. 3a exemplary sagittal and coronal 3D ufSSFP chest imagesillustrating end-expiration (top) and full inspiration (bottom);

FIG. 3b the MR images of FIG. 3a after median-filtering and withsegmentation (white outline) used for estimation of parenchymal signalintensity and lung volume;

FIG. 4 the mean signal intensity of the parenchyma (dots) as a functionof the lung volume in double logarithmic scale as well as a fit of theadapted sponge model (black line) yielding α_(global)=1.16±0.02;

FIG. 5a another exemplary 3D ufSSFP coronal view with threeregion-of-interests 1, 2 and 3 being indicated by respective numbers;

FIG. 5b the local parenchymal signal (averaged over 4×4×4 voxels) as afunction of the lung volume in double logarithmic scale for the threeregion-of-interests indicated in FIG. 5a , yielding α₁=0.81±0.07,α₂=0.93±0.07 and α₃=1.13±0.05;

FIG. 6 exemplary 3D respiratory α-maps overlaid on 3D ufSSFPmorphological images in sagittal, coronal and transverse orientation fora healthy volunteer;

FIG. 7a exemplary CT emphysema maps in sagittal, coronal and transverseorientation for a patient suffering under chronic obstructive pulmonarydisease (COPD);

FIG. 7b images acquired by means of dynamic contrast-enhanced MRI(DCE-MRI) in the same views and in the same patient as in FIG. 7a , witharrows indicating functional defects; and

FIG. 7c respiratory index maps obtained by means of the inventive methodin the same views and in the same patient as in FIG. 7a , with arrowsindicating functional defects.

DESCRIPTION OF PREFERRED EMBODIMENTS

In FIG. 1, an exemplary MRI system is shown which serves to carry outthe inventive method for the assessment of the lung parenchyma in ahuman or an animal. A flow chart illustrating an example of an inventivemethod is shown in FIG. 2.

The MRI system comprises a main magnet 1 for producing a main magneticfield B₀. The main magnet 1 usually has the essential shape of a hollowcylinder with a horizontal bore. Inside the bore of the main magnet 1 amagnetic field is present, which is essentially uniform at least in theregion of the isocenter 6 of the main magnet 1. The main magnet 1 servesto at least partly align the nuclear spins of a sample 5 arranged in thebore. The sample 5 is represented by a human or an animal. Of course,the magnet 1 does not necessarily be cylinder-shaped, but could forexample also be C-shaped.

The human or animal patient is arranged in such a way on a moving table4 in the bore of the main magnet 1, that the lung, the function of whichis being assessed, is arranged in the region of the isocenter 6 of themagnet 1. In order to avoid motion artefacts, the patient is instructedto not move during data acquisition.

The main magnet 1 has a z-axis 9 which coincides with the centrallongitudinal axis defined by the cylindrical shape of the magnet 1.Together with a x-axis 7 and a y-axis 8, which each extend in mutuallyperpendicular directions with respect to the z-axis 9, the z-axis 9defines a Cartesian coordinate system of the MRI system, having itsorigin at the isocenter of the magnet 1.

In order to produce a magnetic field which linearly varies in thedirection of the x-axis 7, the y-axis 8 and/or the z-axis 9, theMRI-system comprises a gradient system 2 including several coils forproducing these varying magnetic fields. A radiofrequency (RF) coil 3 isprovided for generating a transmit field B₁, in order to repetitivelyexcite the nuclear spins of the sample 5.

As illustrated in FIG. 2, the sample 5 is subjected to a 3D ultra-fastbalanced SSFP (ufSSFP) sequence. The 3D ufSSFP sequence is initiated byan operator by means of a user interface 10, which sends the respectiveinstructions to a central control unit of the MRI system (FIG. 1). Thecentral control unit controls a gradient field control unit beingconnected with the gradient system 2 as well as a RF transmitter and areceiver both being connected with the RF coil 3. The gradient system 2and the gradient field control unit together constitute a gradientmodule of the MRI system. The central control unit controls the gradientfield control unit, the RF transmitter and the receiver such, that a 3DufSSFP imaging sequence is employed on the MRI system, which allows theacquisition of MR images of the entire lung parenchyma at differentbreathing positions. The patient is instructed to hold his breath duringthe data acquisition accordingly. Of course it is also possible toacquire the required MR images in free breathing.

The receiver, which constitutes an acquisition module together with theRF coil 3, is connected with a reconstruction unit, in which theacquired MR signals are reconstructed to form the respective MR imagedata sets. After completed image acquisition, these image data sets,which are still spatial spectral image data sets in k-space aretransferred to image space by means of an (inverse) Fouriertransformation carried out by the reconstruction unit (imagereconstruction in FIG. 2). In order to account for the specific imagingparameters applied during image acquisition, the reconstruction unit isconnected with the central control unit. The reconstructed MR images arethen sent to the user interface 10.

In a concrete measurement, all datasets, i.e. MR images, were acquiredwith a 1.5 T clinical whole-body MR-scanner (Siemens Healthcare,Erlangen, Germany) using a 12-channel thorax and a 24-channel spine coilreceive-array and the body coil as the transmitter. Images of humanpatients were acquired in breath-hold and in supine position using a 3DufSSFP protocol, with the following imaging parameters: TE/TR=0.47/1.19ms, flip angle 23°, RF pulse length=80 μs, 1563 Hz/pixel bandwidth,field-of-view=450²×200-250 mm³, two averages, isotropic resolution=3.5mm³, reconstruction matrix=128²×72-80, parallel imaging with a GRAPPAfactor 2, phase oversampling 50%, for a total acquisition time=11-17 s.Predefined default shim settings (tune MR up) were used. Five to 14scans, i.e. MR images, were recorded at different breathing positionsfrom forced expiration to forced inspiration.

In the user interface 10, which also represents an analysis module andis usually realized by a customary personal computer having a processor,the acquired and reconstructed MR images are further processed, in orderto obtain respiratory α-maps of the lung parenchyma.

In order to obtain the respiratory α-maps, the signal intensitiesSI({right arrow over (x)}_(i)) of a plurality of positions or regions{right arrow over (x)}_(i) of the lung parenchyma need to be determinedfor the different breathing positions. In parallel to the determinationof the signal intensities, the lung volumes V_(L) are estimated for eachof the different breathing positions.

As shown in FIG. 2, for the determination of the signal intensitiesSI({right arrow over (x)}_(i)), in a first step the reconstructed MRimages are spatially co-registered in the area of the lung parenchyma.The spatial co-registration can be achieved using a 3D deformableB-spline mass preserving image registration algorithm. This can forexample be done using the open source software elastix (elastix version4.7, University Medical Center Utrecht, The Netherlands). Forregistration, a scan with a mean value of the lung volume can be chosenas fixed dataset (baseline) and the other, so-called ‘moving’ datasetscan be co-registered.

The co-registered MR images are then median filtered to remove theoverlying vascular structure, but preserving the underlying pulmonarysignal variation. With the data sets acquired with the MR imagingparameters as set out above, it is proposed to use a filter kernel sizeof 5×5×5 voxels for median filtering.

From the spatially co-registered and median filtered images, the signalintensities SI({right arrow over (x)}_(i)) can then be determined ineach of the MR images acquired at different breathing positions for aplurality of positions or regions {right arrow over (x)}_(i) in astraightforward way.

The estimation of the lung volumes V_(L) in the different breathingpositions is preferably based on the same acquired MR images as thedetermination of the signal intensities SI({right arrow over (x)}_(i)).As shown in FIG. 2, in a first step the reconstructed MR images aremedian filtered to remove the overlying vascular structure, butpreserving the underlying pulmonary signal variation. As in thedetermination of the signal intensities SI({right arrow over (x)}_(i)),it is proposed to use a filter kernel size of 5×5×5 voxels for medianfiltering the data sets acquired with the MR imaging parameters as setout above.

The lung volumes are then segmented in each of the median filtered MRimages. For the segmentation, it is proposed to use a 2D or 3Dfast-marching algorithm and to verify the consistency of thesegmentation slice-by-slice and in between datasets by a trainedobserver. If needed, the regions-of-interest (ROI) can be correctedmanually.

Based on the segmentation, the volumes V_(L) of the lung in eachacquisition can be calculated by the analysis module.

Finally, the respiratory indices α({right arrow over (x)}_(i)) aredetermined by the analysis module according to the formula

${{\alpha \left( {\overset{\rightarrow}{x}}_{i} \right)} = {{- \frac{1}{1 - {SNR}^{- 1}}} \cdot \frac{d\left( {\log \left( {{SI}\left( {\overset{\rightarrow}{x}}_{i} \right)} \right)} \right)}{d\left( {\log \left( V_{L} \right)} \right)}}},$

SNR denotes an estimated signal-to-noise ratio SI({right arrow over(x)}_(i))/η in which η represents the estimated intensity of noise inthe MR images in the regions or positions {right arrow over (x)}_(i). Byapplying the optional factor

$\frac{1}{1 - {SNR}^{- 1}},$

more reliable α-values can be obtained, which allow for a quantitativeassessment of regional lung function independently of the SNR.

The noise η is preferably determined in the MR images by a ROI analysis,e.g. a ROI placed in the air and away from the body. In this case, thesame SNR is assumed for all MR images and for all regions or positions{right arrow over (x)}_(i). Alternatively, the same MR sequence as usedfor the 3D ufSSFP acquisition can be run with flip angles of 0°, inorder to assess the noise η in the images.

The determination of the respiratory indices α({right arrow over(x)}_(i)) is carried out by means of a linear fit in the doublelogarithmic scale of SI({right arrow over (x)}_(i)) and V_(L) (see FIG.4, as explained further below). The obtained respiratory indicesα({right arrow over (x)}_(i)) are then graphically represented on arespiratory α-map, in which the values of the respiratory indicesα({right arrow over (x)}_(i)) are preferably color-coded and overlaid ona respective 3D ufSSFP morphological image.

Exemplary native 3D ufSSFP images obtained in a healthy volunteer areshown in FIG. 3a for two different inspiratory volumes (i.e.end-expiration with 1.8 L in the top row and full inspiration with 4.6 Lin the bottom row). As a result of the very short TR, no bandingartefacts are visible and airways are well visualized in the nativeimages. The difference in signal levels in the lung parenchyma for thetwo lung volumes are clearly noticeable; especially in the dependent(dorsal) part of the lung. Median filtering, as shown in FIG. 3b , isused to remove the overlying vascular structure, but preserving theunderlying pulmonary signal variation (24). The segmentation of the lungis shown in FIG. 3b by means of the white outline.

The observed mean parenchymal signal intensity variation (over the wholelung; see white outlines in FIG. 3b for definition of theregion-of-interest) is shown in FIG. 4 as a function of lung volumesV_(L) for the MR images acquired with the parameters as set out above.The overall SI variation of the lung parenchyma is as low as 36 [a.u.]in forced inspiration (V_(L)=6.31 L) and increased about 4-times, up to147 [a.u.], in forced expiration (V_(L)=1.92 L). In comparison to thesignal intensity emanating from the lung parenchyma, a low noise levelof about 5±2 [a.u.] is measured in air justifying the assumption ofη<<Λ*V_(L) ^(−α).

Overall, the signal variation as shown in FIG. 4 almost perfectlyfollows the adapted sponge model. A linear fit of the adapted spongemodel in the double logarithmic scale yields α_(global)=1.16±0.02 incombination with a SI scaling factor Λ=317±7 [a.u.]. The small residualdeviations of the experimental data from the linear fit in the doublelogarithmic scale corroborate the overall independency of therespiratory index from the breathing phase. Moreover, a respiratoryindex α_(global) deviating from 1 indicates a change in lung mass duringbreathing and that the dry sponge model (α=1) is not accurate todescribe the lung SI modulations.

Another exemplarily coronal view of a 3D ufSSFP acquisition in a healthyvolunteer is shown in FIG. 5a . Four other images which are not shownwere acquired at different breathing positions. From this imageacquisition, a global respiratory index α_(global)=1.04±0.03 isrecovered.

For three different regions-of-interests (4×4×4 voxels), the mean SI, asobserved in the five datasets after image co-registration, is shown inFIG. 5b . Local respiratory index values α_(local) of 0.88±0.07,0.93±0.07 and 1.13±0.05 are found with an R² exceeding 0.976, indicatingappropriate description of the data by the adapted sponge model afterimage registration.

Respiratory index maps for a healthy male volunteer are shown in FIG. 6for representative slices in coronal, sagittal, and axial orientations.Generally, a homogenous respiratory index is observed onisogravitational planes, whereas the respiratory index values almostdouble in the caudal part (α≈1.4) in comparison to ventral part (α≈0.7).

Respiratory index maps obtained in a patient suffering under chronicobstructive pulmonary disease (COPD) are shown in FIG. 7c and arecompared to CT (FIG. 7a ) and DCE MRI (FIG. 7b ). CT emphysema mapsreveal impairments for the COPD patient: There are localized largeemphysematous regions. Similar abnormalities are seen on DCE perfusionimages (arrows in FIG. 7b ). Overall, structural emphysematous tissuedestructions and perfusion defects visually correlate with the locationof conspicuous functional impairment indicated on the respiratory maps(e.g. α_(local)<0.3; arrows in FIG. 7c ). FIGS. 7 a, b, c demonstratethe feasibility of respiratory index mapping by means of the inventivemethod in the clinical setting.

REFERENCE NUMERALS

1 Main magnet 2 Gradient system 3 RF coil 4 Moving table 5 Patient 6Isocenter 7 X-axis 8 Y-axis 9 Z-axis 10 User interface

1. A method for the assessment of the lung parenchyma in a human or ananimal using a series of magnetic resonance (MR) images of the lungparenchyma acquired at different breathing positions in the same humanor animal, comprising at least the following steps: a.) estimation of achange of a lung volume V_(L) between the different breathing positions;b.) determination of a signal intensity SI({right arrow over (x)}_(i))in at least one same region or position {right arrow over (x)}_(i) ofthe lung parenchyma for each of the MR images; and c.) determination ofat least one respiratory index α({right arrow over (x)}_(i)) accordingto the formula${\alpha \left( {\overset{\rightarrow}{x}}_{i} \right)} = {- {\frac{d\left( {\log \left( {{SI}\left( {\overset{\rightarrow}{x}}_{i} \right)} \right)} \right)}{d\left( {\log \left( V_{L} \right)} \right)}.}}$2. The method as claimed in claim 1, wherein the acquired MR images arespatially co-registered in the area of the lung parenchyma, and wherein,in step b.), the signal intensity SI({right arrow over (x)}_(i)) isdetermined for each of the spatially co-registered MR images.
 3. Themethod as claimed in claim 2, wherein the spatial co-registration of theMR images is carried out by means of a deformable B-spline masspreserving image registration algorithm.
 4. The method as claimed inclaim 2, wherein the spatially co-registered MR images aremedian-filtered, and wherein, in step b.), the signal intensitySI({right arrow over (x)}_(i)) is determined for each of the spatiallyco-registered and median-filtered MR images.
 5. The method as claimed inclaim 1, wherein the estimation of the lung volume V_(L) for each of thedifferent breathing positions is based on the acquired series of MRimages.
 6. The method as claimed in claim 5, wherein each of theacquired MR images is segmented with respect to the lung parenchyma, andwherein, in step a.), the lung volume V_(L) is estimated for each of thedifferent breathing positions based on the segmented MR images.
 7. Themethod as claimed in claim 6, wherein the acquired MR images aremedian-filtered prior to the segmentation.
 8. The method as claimed inclaim 1, wherein the series of MR images comprises at least threeimages, and wherein, in step c.), the at least one respiratory indexα({right arrow over (x)}_(i)) is determined by means of a linear fit. 9.The method as claimed in claim 1, wherein steps b.) and c.) are carriedout for a plurality of regions or positions {right arrow over (x)}_(i)of the lung parenchyma.
 10. The method as claimed in claim 9, whereinthe determined respiratory indices α({right arrow over (x)}_(i)) areused to generate a two-dimensional or three-dimensional map.
 11. Themethod as claimed in claim 1, wherein, in the determination of the atleast one respiratory index α({right arrow over (x)}_(i)) in step c.), amultiplication with the factor $\frac{1}{1 - {SNR}^{- 1}}$ is furthercarried out, in order to yield the formula${{\alpha \left( {\overset{\rightarrow}{x}}_{i} \right)} = {{- \frac{1}{1 - {SNR}^{- 1}}} \cdot \frac{d\left( {\log \left( {{SI}\left( {\overset{\rightarrow}{x}}_{i} \right)} \right)} \right)}{d\left( {\log \left( V_{L} \right)} \right)}}},$wherein SNR denotes an estimated signal-to-noise ratio SI({right arrowover (x)}_(i))/η in which η represents the estimated intensity of noisein the MR images in the at least one region or position {right arrowover (x)}_(i).
 12. The method as claimed in claim 1, additionallycomprising the following step: a0.) acquiring a series of MR images ofthe lung parenchyma at different breathing positions by means of a MRimaging system.
 13. The method as claimed in claim 12, wherein, in stepa0.), the series of MR images is acquired by means of a balanced steadystate free precession (SSFP) imaging sequence.
 14. A computer programfor the assessment of the lung parenchyma in a human or an animal bymeans of a series of magnetic resonance (MR) images of the lungparenchyma acquired at different breathing positions in the same humanor animal, the computer program at least comprising executableinstructions to: b.) determine a signal intensity SI({right arrow over(x)}_(i)) in at least one same region or position {right arrow over(x_(i))} of the lung parenchyma for each of the MR images; and c.)determine at least one respiratory index α({right arrow over (x)}_(i))according to the formula${{\alpha \left( {\overset{\rightarrow}{x}}_{i} \right)} = {- \frac{d\left( {\log \left( {{SI}\left( {\overset{\rightarrow}{x}}_{i} \right)} \right)} \right)}{d\left( {\log \left( V_{L} \right)} \right)}}},$wherein V_(L) denotes a lung volume for each of the different breathingpositions.
 15. The computer program as claimed in claim 14, additionallycomprising executable instructions to a0.) acquire a series of MR imagesof the lung parenchyma at different breathing positions by means of a MRimaging system; and a.) estimate a change of the lung volume V_(L)between the different breathing positions based on the MR images. 16.The computer program as claimed in claim 14, wherein, in thedetermination of the at least one respiratory index α({right arrow over(x)}_(i)) in step c.), a multiplication with the factor$\frac{1}{1 - {SNR}^{- 1}}$ is further carried out, in order to yieldthe formula${{\alpha \left( {\overset{\rightarrow}{x}}_{i} \right)} = {{- \frac{1}{1 - {SNR}^{- 1}}} \cdot \frac{d\left( {\log \left( {{SI}\left( {\overset{\rightarrow}{x}}_{i} \right)} \right)} \right)}{d\left( {\log \left( V_{L} \right)} \right)}}},$wherein SNR denotes an estimated signal-to-noise ratio SI({right arrowover (x)}_(i))/η in which η represents the estimated intensity of noisein the MR images in the at least one region or position {right arrowover (x)}_(i).
 17. A magnetic resonance imaging (MRI) system at leastcomprising a main magnet for generating a main magnetic field at alocation of a sample to be imaged, in order to at least partly alignnuclear spins of the sample; an excitation module for applying radiofrequency (RF) pulses to the sample, in order to repeatedly excite thenuclear spins of the sample; a gradient module for generating temporarymagnetic gradient fields at a location of the sample; an acquisitionmodule for acquiring the magnetic resonance signals produced by excitednuclear spins of the sample; a control module configured for controllingthe excitation module, the gradient module and the acquisition modulesuch, that a series of MR images of the lung parenchyma of a human or ananimal can be acquired at different breathing positions; and an analysismodule being configured for determining a signal intensity SI({rightarrow over (x)}_(i)) in at least one same region or position {rightarrow over (x)}_(i) of the lung parenchyma for each of the MR images andbeing configured for determining at least one respiratory index α({rightarrow over (x)}_(i)) according to the formula${{\alpha \left( {\overset{\rightarrow}{x}}_{i} \right)} = {- \frac{d\left( {\log \left( {{SI}\left( {\overset{\rightarrow}{x}}_{i} \right)} \right)} \right)}{d\left( {\log \left( V_{L} \right)} \right)}}},$wherein V_(L) denotes a lung volume for each of the different breathingpositions.
 18. The MRI system as claimed in claim 17, the analysismodule further being configured for estimating a change of the lungvolume V_(L) between the different breathing positions based on the MRimages.
 19. The MRI system as claimed in claim 17, wherein the analysismodule is further configured for carrying out a multiplication with thefactor $\frac{1}{1 - {SNR}^{- 1}}$ in the determination of the at leastone respiratory index α({right arrow over (x)}_(i)), in order to yieldthe formula${{\alpha \left( {\overset{\rightarrow}{x}}_{i} \right)} = {{- \frac{1}{1 - {SNR}^{- 1}}} \cdot \frac{d\left( {\log \left( {{SI}\left( {\overset{\rightarrow}{x}}_{i} \right)} \right)} \right)}{d\left( {\log \left( V_{L} \right)} \right)}}},$wherein SNR denotes an estimated signal-to-noise ratio SI({right arrowover (x)}_(i))/η in which η represents the estimated intensity of noisein the MR images in the at least one region or position {right arrowover (x)}_(i).